The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to the accurate generation of cardiac gating signals for use in MR imaging and spectroscopy.
The data required to reconstruct an MR image is acquired by an MRI system over a period of time. In most acquisitions this time period extends over many cardiac cycles of the patient and sometimes it is necessary to synchronize the acquisition with the cardiac cycle. This is accomplished in most applications by monitoring an ECG signal produced by the patient's heart and triggering, or gating, the data acquisition sequence when the R-peak in the QRS complex is detected.
The accurate detection of the R wave peak in the ECG signal is very difficult in an MRI system environment. First, the quality of the ECG signal itself is seriously degraded by the magnetic induction effects caused by the strong magnetic fields used in MRI systems. Significant inductive noise is added to the ECG signal by patient movement, heart motion, and blood flow as well as “gradient noise” produced by the rapidly changing magnetic field gradients used during MRI acquisitions. Under the best conditions the production of a reliable ECG trigger signal is very challenging.
In a large number of patients the ECG gating does not work well. In patients who are large with a lot of subcutaneous fat, who have chronic obstructive pulmonary disease with expanded lungs, or for other reasons, the ECG signal may be small and difficult to detect within the electrically noisy environment of an MR imaging system. In addition, the magnetic field influences the shape of the ECG signal. This change in the shape of the ECG signal may make it difficult to correctly synchronize off the QRS complex. This is because the T wave may be increased in size and become difficult to distinguish from the QRS complex.
A solution to this problem has been to use a pulse oximeter such as that disclosed in U.S. Pat. No. 5,743,263 to gate from detected flow related changes in the finger tip. However, the detection of systole at the finger tip is delayed compared to systole in the ventricles. The timing between the two will depend on multiple factors including vascular tree compliance, vascular tree resistance, cardiac output, distance from the heart to name only a few. As a result, this often is not useful for “freezing” the motion of the heart.
In addition to cardiac imaging, there are other situations in which cardiac gating is required. For flow imaging in peripheral vessels, frequently one wants to image these vessels when they contain the maximum amount of flow, i.e. during systole. In this situation ECG is problematic because of the time delay problem discussed above. That is, the timing of systole at a peripheral vessel may be very different than systole at the heart.
Another modality for producing images uses ultrasound. There are a number of modes in which ultrasound can be used to produce images of objects. In the so-called “A-scan” method, an ultrasound pulse is directed into the object by an acoustic transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the strength of scatterers in the object and the time delay is proportional to the range of the scatterers from the transducer. In the so-called “B-scan” method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-scan method and their amplitude is used to modulate the brightness of pixels on a display at the time delay. With the B-scan method, enough data are acquired from which a 2D image of the scatterers can be reconstructed. Another way to represent ultrasound information is called “M-mode”. In this technique, a single B-mode line is reproduced repeatedly and displayed as a time plot with the ordinate corresponding to position along the line and the abscissa representing time. If there is motion along the line, the pixel brightnesses are modulated as structures move in and out of or along the line of sight. Thus moving objects are well represented in this mode.
Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements sandwiched between a pair of electrodes. Such piezoelectric elements are typically constructed of lead zirconate titanate (PZT), polyvinylidene diflouride (PVDF), or PZT ceramic/polymer composite. The electrodes are connected to a voltage source, and when a voltage is applied, the piezoelectric elements change in size at a frequency corresponding to that of the applied voltage. When a voltage waveform is applied, the piezoelectric element emits an ultrasonic wave into the media to which it is coupled at the frequencies contained in the excitation waveform. Conversely, when an ultrasonic wave strikes the piezoelectric element, the element produces a corresponding voltage across its electrodes. Typically, the front of the element is covered with an acoustic matching layer that improves the coupling with the media in which the ultrasonic waves propagate. In addition, a backing material is coupled to the rear of the piezoelectric element to absorb ultrasonic waves that emerge from the back side of the element so that they do not interfere. A number of such ultrasonic transducer constructions are disclosed in U.S. Pat. Nos. 4,217,684; 4,425,525; 4,441,503; 4,470,305 and 4,569,231.
When used for ultrasound imaging, the transducer often has a number of piezoelectric elements arranged in an array and driven with separate voltages (apodizing). By controlling the time delay (or phase) and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (transmission mode) combine to produce a net ultrasonic wave that travels along a preferred beam direction and is focused at a selected point along the beam. By controlling the time delay and amplitude of the applied voltages, the beam with its focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (receiver mode). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the signal from each transducer array element.
This form of ultrasonic imaging is referred to as “phased array sector scanning”. Such a scan is comprised of a series of measurements in which the steered ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, the transmission and reception are steered in the same direction (θ) during each measurement to acquire data from a series of points along an acoustic beam or scan line. The receiver is dynamically focused at a succession of ranges (R) along the scan line as the reflected ultrasonic waves are received. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio.
Cardiac gating is also used when performing ultrasonic imaging. As disclosed in U.S. Pat. No. 5,709,210, for example, when cardiac gating is required a standard ECG signal is used to trigger the image acquisition.